Digital imaging systems are becoming increasingly widespread for producing digital data which can be reconstructed into useful radiographic images. In one application of a digital imaging system, radiation from a source is directed toward a subject, typically a patient in a medical diagnostic application, and a portion of the radiation passes through the subject and impacts a detector. The surface of the detector converts the radiation to light photons, which are sensed.
The detector is divided into an array of discrete picture elements or pixels, and encodes output signals based upon the quantity or intensity of the radiation impacting each pixel region. Because the radiation intensity is altered as the radiation passes through the subject, the images reconstructed based upon the output signals may provide a projection of tissues and other features similar to those available through conventional photographic film techniques.
In available digital detectors, the detector surface is divided into an array of picture elements or pixels, with rows and columns of pixels being organized adjacent to one another to form the overall image area. When the detector is exposed to radiation, photons impact a scintillator coextensive with the image area. A series of detector elements are formed at row and column crossing points, each crossing point corresponding to a pixel making up the image matrix. In one type of detector, each element consists of a photodiode and a thin film transistor. The transistors and photodiodes are typically constructed of amorphous silicon, over which cesium iodide is deposited. The cesium iodide absorbs the X-rays and converts them to light, which is then detected by the photodiodes. As can be appreciated by those skilled in the art, the photodiode acts as a capacitor, where the photodiode is generally charged to some known voltage. Light that is emitted in proportion to the X-ray flux then partially discharges the photodiode. Once the exposure is completed, charge on the photodiode is restored to the initial charge. The amount of charge required to restore the initial voltage on the photodiode is then measured which becomes a measure of the X-ray radiation impacting the pixel during the length of the exposure.
In a detector of the type described above, the cathode of each photodiode is connected to the source of a transistor, and the anodes of all diodes are connected to a negative bias voltage. Further, the gates of the transistors in a row are connected together and a row electrode is connected to scanning electronics. The drains of the transistors in each column are connected together and each column electrode is connected to additional readout electronics. Sequential scanning of the rows and simultaneous read out of the signals present at the column electrode permits the system to acquire the entire array or matrix of signals for subsequent signal processing and display. Thus, the detector is read or “scrubbed” on a row-by-row basis as controlled by the transistor associated with each photodiode. It should be noted that reading of the detector refers to recharging of the photodiodes and collection of data when the image produced by the detector contains valuable data, mainly images that contain exposure or offset data, whereas “scrubbing” refers to similar recharging even though the data is not important and can be discarded or not collected at all. Scrubbing, therefore is executed to restore the charge on the photodiode. More importantly, scrubbing is performed to maintain a proper bias on the diodes during idle periods or to reduce the effects of lag, which is an incomplete charge restoration of the photodiodes. Thus, scrubbing is typically performed to restore and maintain the charge of the photodiode, in addition to preventing a continuous DC voltage bias on the amorphous silicon transistors.
The detector is typically read or scrubbed according to its array structure, that is, on a row-by-row basis. As mentioned above, the reading or scrubbing of the detector is controlled by field effect transistors (FETs) associated with each photodiode. The FET allows the minimization of the number of electrical contacts that would need to made to the detector. Thus, the FETs reduce the number of required contacts to no more than the number of pixels along the perimeter of the array. Further, an entire row of an array may be controlled simultaneously when the scan line attached to the gates of all the FETs of pixels on that particular row is activated. It should be noted that each of the pixels on the particular row is connected to a separate data line through a FET, which is used by the readout electronics to restore the charge to the photodiodes. Thus, as each row is activated or read-enabled, the charge is restored for each pixel in that row simultaneously by the read out electronics over the individual data line for each column. Thus, each data line has a dedicated read out channel associated with it.
In use, the signals generated at the pixel locations of the detector are sampled and digitized. The digital values are transmitted to processing circuitry where they are filtered, scaled, and further processed to produce the image data set. The data set may then be used to reconstruct the resulting image, to display the image, such as on a computer monitor, to transfer the image to conventional photographic film, and so forth. In the medical imaging field, such images are used by attending physicians and radiologists in evaluating the physical conditions of a patient and diagnosing disease and trauma.
Large area solid-state detector arrays, such as those described above, provide solutions for digital imaging applications such as medical imaging, digital reproduction and non-destructive testing. While such detectors provide excellent image data, further improvement is needed. For instance, a detector having a small pixel size providing high spatial resolution comes at the cost of acquisition frame rate. Furthermore, for a given pixel size, a larger detector is generally more expensive due to the cost of the required support electronics, and will not support frame rates as fast as a smaller detector with the same bandwidth. Many times, when a specific application is targeted, detector design tradeoffs are made to optimize the detector's performance in regards to that application. One result of this is that larger detectors with small pixels typically do not have the bandwidth to support higher frame rates.
There is a need, therefore, for a technique designed to increase the acquisition frame rate without affecting other variables that reduce image quality.